Composite viscoelastic hydrogel, and uses thereof for sealing a channel in tissue

ABSTRACT

A composite viscoelastic hydrogel comprises a continuous phase of non-crosslinked hyaluronic acid gel and a dispersed phase of dehydrothermally-crosslinked micron-sized gelatin hydrogel particles. The hydrogel exhibits a storage modulus (G′) of greater than 400 Pa and a tan δ (G″/G′) from 0.1 to 0.8 in dynamic viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate at 25° C. The dehydrothermally crosslinked micron-sized gelatin hydrogel particles have an average dimension of less than 100 microns prior to hydration.

FIELD OF THE INVENTION

The present invention relates to a composite viscoelastic hydrogel, and uses thereof for sealing a channel in tissue. The invention also relates to methods, kits and systems employing a composite viscoelastic hydrogel, including methods, kits and systems for sealing a channel in tissue.

BACKGROUND TO THE INVENTION

A number of surgical procedures require puncturing an instrument through the body to gain access to a target treatment region, such as puncturing the thoracic wall to gain access to the thoracic cavity. The most common example is transthoracic needle lung biopsy where a special needle is used to obtain a sample of tissue from a suspected cancerous tissue mass. This procedure, which is presented schematically in FIGS. 1A-1D (Prior art), is typically carried out by an interventional radiologist using CT (computed tomography) guidance. When the biopsy needle punctures the outer surface of the lung air can escape between the lung and the thoracic wall into a space known as the pleural cavity. The air gradually pushes the lung away from the thoracic wall causing the lung to collapse, a complication known as pneumothorax. If the pneumothorax is large, it can lead to severe pain and distress for the patient. An unresolved pneumothorax can lead to the patient being admitted to hospital for treatment and monitoring and often requires the surgical insertion of a chest drain to withdraw the air in the pleural cavity. Pneumothorax can result in considerable pain and morbidity to the patient, increased anxiety and stress to the attending clinician, and unnecessary and substantial costs to the hospital. Approximately 33% of patients undergoing a transthoracic lung biopsy procedure will develop a pneumothorax and approximately 1 in 3 of these patients will require a chest drain.

Methods to prevent pneumothorax are of great interest because of the concomitant morbidity and hospital expenditures. Numerous attempts have been described in scientific literature and have focused on plugging the biopsy needle tract with an adhesive or plug as the biopsy needle is being withdrawn. A number of different substances have been injected with this purpose including gelatine sponge slurry, fibrin adhesive, autologous blood, supernatant serum and autologous blood mixture, and collagen foam. These efforts have proven ineffective and have not been widely adopted. Their lack of efficacy may be as a result of the physical properties of the substances injected and the lack of control over their injected location. Additional references which may be suitable for lung sealing are outlined in U.S. Pat. No. 6,592,608B2 and U.S. Pat. No. 6,790,185B1. This technology is commercially available as the Biosentry™ device from Surgical Specialties Corp (MA, USA www.biosentrysystem.com). Other publications relevant to lung and tissue sealing include US2016120528A, US2006025815A, US2013338636A, US2006009801A, U.S. Pat. No. 6,770,070B, US2017232138A, US2002032463A, and US2009136589A.

SUMMARY OF THE INVENTION

The Applicant has discovered that compositions for generating a sealing plug in tissue should be shear-thinning (which allows the viscosity of the composition to drop under shear during injection, and then increase after delivery to form the sealing plug), should be sufficiently robust to prevent the composition separating during injection (i.e. prevent the water being squeezed out of the composition), and requires rheological properties that allow the composition to self-heal after needle delivery, and to exhibit tissue apposing properties (i.e. push tissue away). In addition to these properties, the composition should be biodegradable but also have sufficient in-vivo persistence time to perform a sealing function for a sufficient period of time (for example in the case on a biopsy needle tract, during the period of healing of the needle tract). The Applicant has a discovered that a composite viscoelastic hydrogel having a continuous (non-crosslinked) hyaluronic acid hydrogel phase, and a dispersed phase comprising crosslinked gelatin or collagen hydrogel particles, is ideal for providing a sealing plug in tissue, especially in lung parenchymal tissue. This hydrogel is shear-thinning, tissue-apposing, and biodegradable over a period of a few weeks which provides adequate time for tissue (i.e. a needle tract to heal). In a particular embodiment, the use of a weak crosslinking method for the collagen or gelatin particles (i.e. dehydrothermal crosslinking) has been found to be ideal for providing a hydrogel gel that persists in tissue during a period of therapy and ultimately biodegrades. In addition, the amount of collagen or gelatin particles in the hydrogel may be varied to vary the tissue persistence time An essential benefit of providing a viscous continuous phase (provided by the hyaluronic acid) is to ensure that there is no phase separation of the dispersed phase during injection. As an illustration, if gelatin hydrogel particles were to be suspended in saline solution as a carrier in a syringe, during injection through a narrow gauge needle the saline solution would be expelled from the needle more easily than the suspended gelatin particles. The gelatin particles would condense behind the needle opening and the saline would pass through, concentration the remaining hydrogel in the syringe. This would alter the uniformity of hydrogel injection, which is a key attribute for several medical applications such as the one described in this patent. The composite nature of the gel has another additional benefit in vivo: the colloidal gelatin provides an excellent framework for cellular infiltration and vascularisation.

According to a first aspect of the present invention, there is provided a composite viscoelastic hydrogel comprising a continuous phase and a dispersed polymer phase.

In one embodiment, the dispersed phase is colloidal polymer, typically polymer hydrogel particles. In one embodiment, the dispersed phase is crosslinked, ideally dehydrothermally crosslinked.

In any embodiment, the dispersed phase polymer is crosslinked by dehydrothermal (DHT) crosslinking. This method avoids adding crosslinking chemicals to the gel. In addition, this method of crosslinking has been found to be ideal for achieving a relatively weak crosslinking of the dispersed polymer particles, required to make the particles biodegradable yet persist in the tissue during a tissue healing period (i.e. a period of weeks). Typically, the DHT crosslinking is performed under vacuum at a temperature greater than 100° C., typically 130° C. to 160° C., for an extended period, for example 24 hours. It will be clear to a person skilled in the art that an equivalent level of weak crosslinking may be achieved by using a lower temperature and longer time, or vice versa. With respect to gelatin, during the DHT treatment, water is extracted from gelatin by condensation reactions, such as esterification or amide link formation, resulting in the formation of intermolecular crosslinks that improve the materials resistance to dissolution. When placed in aqueous suspensions, the DHT treated gelatin powders swell by absorbing the surrounding water to form a colloidal hydrogel. The amount of swelling is dependent on the level of crosslinking. This feature greatly influences the final rheological properties of the gel. An additional benefit of employing the DHT crosslinking process in relation to gelatin is that it removes residual endotoxins which may be present in the polymer from processing as the polymer is derived from an animal source (collagen).

In one embodiment, the continuous phase is a hydrogel. In one embodiment, the continuous phase polymer is a glycosaminoglycan (or a salt thereof), for example hyaluronic acid.

In one embodiment, the composite viscoelastic hydrogel comprises 2-25%, 10-20%, 5-15%, 12-18%, 8-12% colloidal polymer.

In one embodiment, the colloidal polymer comprises or consists of gelatin or collagen (or a mixture thereof).

In one embodiment, the composite viscoelastic hydrogel comprises about 1-6%, 2-6%, 0.5-2.0%, or 0.6-1.2% continuous phase polymer (i.e. HA).

In one embodiment, the continuous phase polymer comprises or consists of HA (or another glycosaminoglycan).

In one embodiment, the continuous phase polymer (i.e. HA) is not cross-linked, or is lightly cross-linked.

In one embodiment, the invention provides a composite viscoelastic hydrogel comprising a continuous polymer phase comprising 2-6% polymer, and a dispersed polymer phase comprising 2-25% colloidal polymer in the form of crosslinked polymer microbeads In one embodiment, the crosslinked polymer particles have an average dimension of less than 100 microns prior to hydration.

In one embodiment, the continuous phase polymer comprises or consists of HA and the colloidal polymer comprises gelatin or collagen.

In one embodiment, the composite viscoelastic bi-phasic hydrogel exhibits a storage modulus (G′) of greater than 400 Pa and a tan δ (G″/G′) from 0.1 to 0.8 in dynamic viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate at 25° C. In one embodiment, the composite viscoelastic bi-phasic hydrogel exhibits an axial compressive stiffness of greater than 800 Pa, as measured using an axial compression testing machine.

In one embodiment, the dispersed phase comprises dehydrothermally crosslinked micron-sized gelatin hydrogel particles.

In one embodiment, the composite viscoelastic hydrogel is formulated to exhibit an in vivo degradation period in the lung tissue of at least 2 weeks.

In one embodiment, the dehydrothermally-crosslinked micron-sized gelatin hydrogel particles are formulated to exhibit an in-vivo degradation period in the lung tissue of less than 2 months.

In one embodiment, the dehydrothermally crosslinked micron-sized gelatin or collagen hydrogel particles have an average dimension of less than 100 microns prior to hydration (i.e. in a dehydrated state).

In one embodiment, the hyaluronic acid has a molecular weight of at least 1 MDa, for example 1-2 MDa.

In one embodiment, the DHT gelatin powder is derived from a high Bloom strength gelatin, for example 200-400, preferably 250-350-, and ideally about 300 gBloom.

In one embodiment, the composite viscoelastic bi-phasic hydrogel is shear-thinning. In one embodiment the composite viscoelastic bi-phasic hydrogel demonstrates a storage modulus G′ of less than 200 Pa at 100% strain as measured by a rheometer under strain control and at a test frequency of 1 Hz over a strain range of 0.01-100%.

In one embodiment, the composite viscoelastic shear-thinning bi-phasic hydrogel comprises:

-   -   a continuous phase of non-crosslinked biodegradable hyaluronic         gel; and     -   a dispersed phase of dehydrothermally-crosslinked micron-sized         gelatin or collagen hydrogel particles,     -   in which the hydrogel exhibits a storage modulus (G′) of greater         than 400 Pa and a tan δ (G″/G′) from 0.1 to 0.8 in dynamic         viscoelasticity measured by a rheometer at 1 Hz and 1% strain         rate at 25° C.

In one embodiment, the composite viscoelastic shear-thinning bi-phasic hydrogel comprises:

-   -   a continuous phase of non-crosslinked biodegradable hyaluronic         gel; and     -   a dispersed phase of dehydrothermally-crosslinked micron-sized         gelatin or collagen hydrogel particles,     -   in which the hydrogel comprises 8-25% gelatin or collagen         hydrogel particles having an average dimension of less than 300         microns in a dehydrated state.

In one embodiment, the composite viscoelastic shear-thinning bi-phasic hydrogel comprises:

-   -   a continuous phase of non-crosslinked biodegradable hyaluronic         gel; and     -   a dispersed phase of dehydrothermally-crosslinked micron-sized         gelatin or collagen hydrogel particles,     -   in which the hydrogel comprises 8-25% gelatin or collagen         hydrogel particles having an average dimension of less than 100         microns prior to hydration, and about 0.4-2.0% of hyaluronic         acid.

In another aspect, the invention provides a system for sealing a channel in tissue (for example a channel created during a minimally invasive percutaneous procedure) comprising:

-   -   a medical device comprising a hydrogel delivery needle (4) with         a tip (5) (generally a piercing tip) and a hydrogel outlet (6),         and     -   a composite viscoelastic hydrogel of the invention.

In one embodiment, the composite viscoelastic hydrogel exhibits a storage modulus (G′) of at least 400 Pa in dynamic viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate at 25° C.

In one embodiment, the composite viscoelastic hydrogel exhibits a tan δ (G″/G′) from 0.05 to 0.8 in dynamic viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate at 25° C.

In one embodiment, the composite viscoelastic hydrogel is configured to exhibit an in-vivo residence time of at least 1, 2 or 3 weeks. This enables the gel to persist in tissue, while the tissue needle tract in the tissue heals. Generally, one week is sufficient, but at least two weeks in-vivo residence time is preferred. Hydrogels formed from, or comprising, crosslinked polymers help with in-vivo residence time. For example, by creating a composite hydrogel containing 4-5% non-crosslinked hyaluronic acid and crosslinked gelatin particles (crosslinked by dehyrothermal treatment) an in-vivo residence time of at least two weeks in a lung needle biopsy tract was achieved.

In one embodiment, the composite viscoelastic hydrogel is configured to exhibit an in vivo residence time of less than 3 months, more preferably less than 2 months, and more preferably less than 1 month. A fast degradation period for the is an advantage in several applications, where a long lasting hydrogel can be confused as a cancerous tumour on CT or MRI scans at follow-up periods following procedures. In addition, a fast degrading hydrogel avoids issues with chronic inflammation of the wound site and allows the tissue to return to normal in line with the natural wound healing process.

The composite viscoelastic hydrogel (hereafter “composite viscoelastic hydrogel” or “viscoelastic hydrogel” or “biphasic hydrogel” or “hydrogel” or “gel”) is generally a tissue apposing hydrogel of sufficient properties that limits its infiltration of tissue so that it pushes the tissue away. In this way the hydrogel can create its own discrete space inside a tissue or organ. To achieve this the properties must be present on entering the target injection site. Typically, the viscoelastic hydrogel exhibits a storage modulus (G′) of at least 400 Pa (e.g. 800-6000 Pa), and a tan δ (G″/G′) from 0.1 to 0.8 in dynamic viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate at 25° C. The hydrogel is configured for injection via a syringe.

For improved tissue opposing properties and to form a uniform plug surrounding the needle, it is also preferable that the viscoelastic hydrogel portrays an axial compressive stiffness of equal to or greater than lung parenchymal tissue, as measured using an axial compression testing machine, for example by using a Zwick universal testing machine with a 5N load cell at a strain rate of 3 mm/min. The viscoelastic hydrogel should preferably have a compressive modulus of greater than 200 Pa, preferably greater than 400 Pa, and more preferably greater than 800 Pa.

Optionally in any embodiment, the injectable composite viscoelastic hydrogel is a shear thinning gel. For example, the viscoelastic hydrogel may be configured to have a low viscosity under higher shear stress or shear rates (i.e. during injection through a needle), and a higher viscosity (under lower shear stresses or shear rates) after removal of shear stress (i.e. once delivered to a target location in the body. This enables these materials to create a singular hydrogel plug at the site of delivery. Materials which possess these properties are outlined in the review articles ‘Shear-thinning hydrogels for biomedical applications’, Soft Matter, (2012) 8, 260, ‘Injectable matrices and scaffolds for drug delivery in tissue engineering’ Adv Drug Deliv Rev (2007) 59, 263-272, and ‘Recent development and biomedical applications of self-healing hydrogels’ Expert Opin Drug Deliv (2017) 23: 1-15. Typically, the shear thinning viscoelastic hydrogel exhibits a storage modulus (G′) of less than 200 Pa, preferably less than 100 Pa in dynamic viscoelasticity at a frequency of 1 Hz and 100% strain.

Optionally, in any embodiment, the composite viscoelastic hydrogel is self-healing. This refers to the hydrogel's ability to spontaneously form new bonds between molecules when old bonds are broken within the material.

Optionally in any embodiment, the colloidal hydrogel is formed by hydrating biocompatible polymer particles which are preferably insoluble in biological fluid. Optionally in any embodiment, the degradation period of the polymer particles is preferably less than 1 year, more preferably less than 6 months, and more preferably less than 2 months. Optionally in any embodiment, the colloidal hydrogel is comprised of a polymer of biological origin, for example gelatin, collagen, fibrin or hyaluronic acid. Optionally in any embodiment, the polymer is crosslinked. Optionally in any embodiment, the colloidal hydrogel comprises about 0.2-30%, 15-28%, or 20-27% hydrogel forming polymer (w/v). Optionally in any embodiment, the colloidal hydrogel exhibits a storage modulus (G′) of greater than 400 Pa, more preferably greater than 800 Pa, more preferably greater than 1000 Pa in dynamic viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate at 25° C.

Optionally in any embodiment, the continuous phase hydrogel may be formed by a hyaluronan hydrogel, and may be present at a concentration of 0.4-6%. Optionally in any embodiment the hyaluronan hydrogel may be non-crosslinked or lightly crosslinked.

Optionally in any embodiment, the colloidal hydrogel may be present at concentrations of 0.2 to 30%, 8 to 25%, 8 to 15%, 8 to 12%, or about 10% hydrogel forming polymer (w/v).

Optionally in any embodiment, the colloidal hydrogel is formed from hydrated polymer particles of <100 μm in average particle size (for example 5-99, 20-80, or 30-80 microns.

Optionally in any embodiment the colloidal hydrogel is insoluble in aqueous solution.

Optionally in any embodiment the colloidal hydrogel is formed from crosslinked polymer particles. Optionally in any embodiment, the colloidal hydrogel is a gelatin hydrogel comprising dehydrothermally (DHT) crosslinked gelatin powders having an average particle size (D₅₀) of about 10-100, 20-50 or 30-40 microns. Optionally in any embodiment, the biphasic hydrogel exhibits a storage modulus (G′) of greater than 400 Pa, more preferably greater than 800 Pa, more preferably greater than 1000 Pa, and a tan δ (G″/G′) from 0.1 to 0.6 in dynamic viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate at 25° C.

Optionally in any embodiment, the biphasic hydrogel portrays an axial compressive stiffness of equal to or greater than lung parenchymal tissue, as measured using an axial compression testing machine Optionally in any embodiment, the composite viscoelastic hydrogel is de-aerated which means it has been removed of air and/or gas or in other words de-gassed.

Optionally in any embodiment, the composite viscoelastic hydrogel comprises a therapeutic agent.

Optionally in any embodiment, the composite viscoelastic hydrogel is biodegradable.

Optionally in any embodiment, the composite viscoelastic hydrogel is comprised of 0.5-6%, preferably 3-5% high molecular weight hyaluronan (w/v). Optionally in any embodiment, the hyaluronan hydrogel may be combined with 0.2 to 30% colloidal hydrogel to form a biphasic hydrogel. Optionally in any embodiment, the colloidal hydrogel may be comprised of hydrogel forming polymer particles. Optionally in any embodiment, the hydrogel forming polymer particles are gelatin particles, collagen particles or hyaluronan particles.

Optionally in any embodiment, the composite viscoelastic hydrogel described herein may be provided in separate components, for example in multiple syringes and the means can be provided to allow mixing of the components prior to injection through the syringe.

Optionally in any embodiment, the system and methods described herein include an initial step of providing the viscoelastic hydrogel as a dehydrated or semi-dehydrated powder, and reconstitution of the powder in a suitable fluid to form the viscoelastic hydrogel.

Optionally in any embodiment, the composite viscoelastic hydrogel is a microporous hydrogel which can be described as hydrogels with interconnected pores that can mechanically collapse and recover reversibly. When the hydrogel is delivered via injection with a needle and syringe, water is squeezed out from the pores, which causes the hydrogel to collapse, allowing it to pass through the needle. Once the hydrogel has left the needle and the mechanical constraint imposed by the needle walls is removed, the hydrogel can recover its original shape almost immediately in the body. These hydrogels generally behave like a foam and can be reversibly compressed at up to 90% strain without any permanent damage to the network.

Optionally in any embodiment, the composite viscoelastic hydrogel is provided in a syringe configured for fluidic connection to a proximal end of the hydrogel delivery needle.

Optionally in any embodiment, the syringe comprises 200 μL to 5000 μL of viscoelastic hydrogel, 200 μL to 2000 μL of viscoelastic hydrogel, or 200 μL to 1000 μL of viscoelastic hydrogel.

Optionally in any embodiment, the hydrogel delivery needle diameter can range from 10-24 gauge, preferably from 16-20 gauge. This is the typical needle size range for lung diagnostic procedures. Larger delivery needles (10-16 gauge) may be employed for other procedures including therapeutic procedures such as lung, live and kidney ablation. Smaller needles greater than 20 gauge or larger than 10 gauge may be used for other medical procedures.

Optionally in any embodiment, the hydrogel outlet is spaced proximal to the piercing tip of the needle. The position of the hydrogel outlet on a side of the needle enables formation of a closed annular sealing plug around the needle, and the viscoelastic properties of the hydrogel allow the annular sealing plug to re-shape upon removal of the device whereby the hole in the middle of the sealing plug is filled in. Optionally in any embodiment, the hydrogel outlet is spaced from preferably 1 to 15 mm or more preferably 3-8 mm, from a piercing tip of the needle.

Optionally in any embodiment, the hydrogel delivery needle comprises a plurality of hydrogel outlets disposed on a side of the needle. The hydrogel outlets may be disposed in a radial fashion around the circumference of the needle. The hydrogel outlets may be circular in profile, in which case their size can range from 0.3-1.5 mm in diameter depending on the diameter of the hydrogel delivery needle. The hydrogel outlets may also take non-circular and elongated profiles.

Optionally in any embodiment, the hydrogel outlet consists of a radiolucent region on the delivery needle where sufficient material has been removed through cutting or erosion process to provide a contrast in radiopacity between the delivery needle and the hydrogel outlet.

Optionally in any embodiment, the coaxial cannula consists of an aperture proximal to its distal tip. This aperture may form a radiolucent region on the coaxial cannula by removing sufficient material about the circumference of the cannula.

Optionally in any embodiment, radiolucent regions of both the delivery needle and coaxial cannula are aligned when the delivery needle and cannula are engaged. This will provide a marking function about this radiolucent region during radiographic guidance and allows the viscoelastic hydrogel to be injected at this location.

Optionally in any embodiment, the hydrogel outlet and coaxial cannula aperture may be created using a laser cut profile or pattern which removes a portion of material from the delivery needle wall to create a pathway through which the hydrogel material can flow to the intended target. Removal of a significant amount of material will provide radiolucency to this portion of the device and will provide visual feedback on the position of the hydrogel outlet under CT guidance or other imaging modality. The radiolucency (less radiopaque) is achieved by removal of a significant amount of material from the needle walls using the laser cut pattern without affecting the structural integrity of the needle. Laser cut profiles comprising circumferential triangles and similar structures to those employed in coronary stents can be employed to maintain structural stability. Alternative material eroding technology may also be employed to create the cut pattern.

Optionally in any embodiment, the medical device comprises an adjustable positioning mechanism configured to limit the advancement depth of the hydrogel delivery needle through the coaxial cannula as indicated by a measurement scale forming part of the medical device, and typically forming part of the positioning mechanism.

Optionally in any embodiment, the positioning mechanism comprises a fixed housing attached to the hydrogel delivery needle, a movable hub mounted to the needle for axial movement along the hydrogel delivery needle relative to the fixed housing and having a distal-most face configured to abut a proximal face of the coaxial cannula luer lock.

Optionally in any embodiment, a visible mark is provided on the delivery needle proximally to the piercing tip.

Optionally in any embodiment, the system further comprises a core needle with penetrating distal tip configured for insertion through the inner lumen of the coaxial cannula and attachment to the coaxial cannula luer lock.

Optionally in any embodiment, the system further comprises a syringe configured for fluidic connection to the hydrogel delivery needle, and in which the viscoelastic hydrogel is provided in the syringe.

In another aspect, the invention provides a method of performing a lung needle biopsy, comprising the steps of:

delivering a composite viscoelastic hydrogel of the invention to a target location in the lung of a patient adjacent the visceral pleura of the lung;

advancing the coaxial cannula distally over the hydrogel injection needle and through the sealing plug;

removal of the hydrogel delivery needle through the cannula;

advancing a biopsy needle through the cannula to a biopsy site within the lung; actuating the biopsy needle to take a sample of lung tissue at the biopsy site;

withdrawing the biopsy needle through the cannula; and

withdrawing the cannula whereby the sealing plug seals the visceral pleura.

Optionally in any embodiment, after the removal of the hydrogel delivery needle and prior to advancement of the biopsy needle, the method includes the steps of insertion of a core needle into the coaxial cannula, advancement of the core needle and coaxial cannula to the biopsy site within the lung, and removal of the core needle.

Optionally in any embodiment, prior to removal of the hydrogel delivery needle, the method includes the steps of advancing the hydrogel delivery needle to the biopsy site within the lung, and then advancing the coaxial cannula over the hydrogel delivery needle to the biopsy site within the lung.

Optionally in any embodiment, the step of advancing the coaxial cannula distally over the hydrogel injection needle to the biopsy site in the lung is guided by the cannula depth guide.

Optionally in any aspect, the invention provides a method of performing a lung needle biopsy procedure comprising the steps of:

injecting a viscoelastic hydrogel (for example, a composite viscoelastic hydrogel of the invention) through a hydrogel delivery needle into the lung adjacent the visceral pleura of the lung to form a sealing plug that embraces the needle and abuts the visceral pleura; advancing a coaxial cannula along the hydrogel delivery needle and through the closed annular sealing plug;

removal of the hydrogel delivery needle through the cannula;

advancing a biopsy needle through the cannula to a target location within the lung;

actuating the biopsy needle to take a sample of lung tissue at the target location;

withdrawing the biopsy needle through the cannula; and

withdrawing the cannula whereby the sealing plug seals the visceral pleura preventing pneumothorax.

In another aspect, the invention provides a method of performing a lung nodule localisation procedure comprising the steps of:

injecting a composite viscoelastic hydrogel of the invention through a hydrogel delivery needle into the lung adjacent the visceral pleura of the lung to form a sealing plug that embraces the needle and abuts the visceral pleura;

advancing a coaxial cannula along the hydrogel delivery needle and through the closed annular sealing plug;

removal of the hydrogel delivery needle through the cannula;

advancing a tissue stain delivery needle through the cannula to a target location within the lung;

actuating the tissue stain needle to take a sample of lung tissue at the target location; withdrawing the tissue stain needle through the cannula; and

withdrawing the cannula whereby the sealing plug seals the visceral pleura preventing pneumothorax.

In another aspect, the invention provides a method comprising delivery of a composite viscoelastic hydrogel of the invention into tissue of a patient, especially a lung of a patient adjacent the visceral pleura of the lung to form a sealing plug wholly within the lung that abuts the visceral pleura.

Optionally in any embodiment, the methods of the invention involve delivering a volume of 100 to 3000 μl of hydrogel. Optionally in any embodiment, the methods involve delivering a volume of 100 to 1000 μl of hydrogel. Optionally in any embodiment, the methods involve delivering a volume of 200 to 900 μl of hydrogel. Optionally in any embodiment, the methods involve delivering a volume of 200 to 500 μl of hydrogel.

Optionally in any embodiment, the composite viscoelastic hydrogel is delivered into the lung through a needle having a piercing tip and a hydrogel outlet disposed on a side of the needle spaced apart from piercing tip.

In another aspect, the invention provides a composite viscoelastic hydrogel of the invention for use in forming a sealing plug in a lung of a patient to prevent pneumothorax during a lung needle biopsy procedure, in which the sealing plug is typically delivered to the lung adjacent and abutting a visceral pleura.

Optionally in any embodiment, the biopsy needle is passed through the sealing plug during the needle biopsy procedure.

Optionally in any embodiment, a coaxial cannula is passed through the sealing plug, and the biopsy needle is passed through the sealing plug via the coaxial needle.

Optionally in any embodiment, the target location in the lung is located 0.2 to 6.0 mm distal of the visceral pleura.

Optionally in any embodiment the target location for delivery of the hydrogel material is into the pleural cavity. In this instance the hydrogel outlet will reside inside or across the pleural cavity.

Optionally in any embodiment, the hydrogel delivery needle may have a hydrogel outlet at the tip of the needle as opposed to the side. It is also possible to have both a hydrogel outlet at the tip of the needle and/or on the side of the needle. The delivery device and system described herein may also provide an effective solution to prevent bleeding during procedures requiring minimally invasive percutaneous access to other organs such as the liver and kidney. These procedures may include diagnosis or treatment of part or all of these organs.

Optionally in any embodiment, the system and/or composite viscoelastic hydrogel described herein can be used to separate tissue during a surgical procedure. This may be required to create a pathway through tissue for an instrument or to protect tissue from unwanted stimuli which as tumour ablation or radiotherapy. For this purpose a greater volume of viscoelastic hydrogel may be delivered, for example 1-25 ml.

Optionally in any embodiment, the system and/or the composite viscoelastic hydrogel described herein can be used as to fill voids in tissue or organs.

Optionally in any embodiment, the system and/or the composite viscoelastic hydrogel described herein can be employed in the prevention of adhesion between adjacent tissues and organs.

Optionally in any embodiment, the system and/or composite viscoelastic hydrogel described herein can be employed as a drug delivery vehicle. The composite viscoelastic hydrogel may be loaded with a drug or any other substance having physiological activity which will slowly diffuse from the hydrogel after its implantation into the body and the diffusion rate can be conveniently controlled by changing the compositional parameters of the hydrogel.

Optionally in any embodiment, the system and/or composite viscoelastic hydrogel described herein can be used as an embolic agent for occlusion of an artery or vein. The composite viscoelastic hydrogel can be deployed into an artery or vein to occlude the flow of blood, either on a temporary or permanent basis. In this manner, the hydrogel can be used to treat venous diseases, for example aneurysm, varicose veins, insufficient veins, dilated veins and ectasias. Thus, in one embodiment, the invention provides a method of occluding a lumen (for example a section of vasculature) in a subject comprising a step of delivering a composite viscoelastic hydrogel according to the invention into the lumen to occlude the lumen.

Optionally in any embodiment, the system and/or composite viscoelastic hydrogel described herein can be employed as a tissue bulking agent or tissue augmenting agent.

Optionally in any embodiment, the devices and components described herein may be created using biocompatible materials including polymers, metals and ceramics. Polymers can include Polyether ether ketone, Polyethylene terephthalate, Nylon, polyimides, polyurethanes, polyesters, Pebax® and copolymers thereof. Metals may include stainless steel, nitinol, titanium and cobalt chrome. The needles and cannula may also comprise fully or partially flexible laser cut sections and braided sections to provide flexibility. The needles and cannula may also be both elongated and flexible such as in catheter type assemblies.

In a preferred embodiment, the compositions of the system, or the system as a whole can be provided sterile for clinical use. The hydrogel filled syringe can be prepared through an aseptic formulation, mixing, filling and packaging process. The hydrogel filling syringe may also be terminally sterilized through a heat or steam sterilization process for e.g., autoclaving. Sterilization of the system can also be performed via sterilization processes known in the field including sterilization by ethylene oxide, hydrogen peroxide, gamma ray and electronic beam.

Optionally in any embodiment, the components of the system can be provided in packaging suitable for sterilization including, but not limited to, a pouch, a blister pack, a bag, a procedure set, a tub, a clamshell, a skin pack, a tray (including lid), a carton, a needle sheath. The components of the system can all be assembled as a single packaged device. Alternatively, multiple packages containing the different components of the system can be prepared and sterilized separately. The components of the system can include but are not limited to the coaxial cannula with core needle, the hydrogel delivery needle, the cannula depth lock, locking arm, one or more syringes filled with viscoelastic hydrogel, empty syringes, hypodermic needles, scalpels, skin markers, radiopaque guides, scissors, biopsy needles, surgical drapes, antiseptic solution, swabs, swab holders, sponges, saline solution and histology tissue containers.

Optionally in any embodiment, the methods described herein include an initial step of flushing the syringe with gel (or saline or water) prior to insertion of the needle into the body. The syringe may also be flushed with the hydrogel prior to insertion into the body.

Optionally in any embodiment, the piercing tip of the delivery needle is designed to prevent bleeding on insertion into the lung, for example it may have a non-cutting atraumatic needle tip profile, for example a pencil tip style needle or similar will help prevent bleeding.

Optionally in any embodiment, the piercing tip is designed with a sharpened bevel profile to minimise disruption of the parietal and visceral pleural layers as the needle is being advance through to the lung.

Optionally in any embodiment, the tip of the delivery needle may be blunt. Optionally in any embodiment the hydrogel outlet may be positioned distal to the blunt tip. Optionally in any embodiment the tip of the delivery needle may be configured with a veress needle tip that combines a spring activated blunt core and a sharp piercing tip.

Optionally in any embodiment the delivery needle is a single lumen. Optionally in any embodiment the delivery needle is comprised of a multi-lumen tube. The multi-lumen tube may be a single tube, or may be comprised of multiple individual tubes within another lumen (for example a stainless steel needle). The tubes may be connected to different delivery outlets. For example, one tube may be connected to a delivery outlet that is distal to the needle tip, whereas the other lumen may be connected directly to the needle tip. Individual delivery lumens may be used to deliver the hydrogel, deliver instruments, take measurements (pressure, temperature, impedance), extract tissue (for example FNA or core biopsies). The tubes may also be used to delivery crosslinking agents, chemotherapy agents and cellular solution (for example stem-cells).

Optionally in any embodiment the delivery needle may be comprised of a single tube.

Optionally the single tube may comprise a tissue penetrating tip. Optionally the delivery needle may be comprised of two or more tubes bonded together, whereby the distal tube may form a tissue penetrating tip. The various tubes used to comprise the delivery needle can be made from radiodensity contrasting materials, for example stainless steel or polymer.

Optionally in any embodiment, the delivery needle can be provided with a central lumen to allow it to pass over a guidewire. The guidewire can be provided for access to body cavities or lumens.

Optionally in any embodiment the delivery needle and coaxial cannula can be given atraumatic and friction prevention properties by use of surface coatings and surface modifications such as polytetrafluorinated ethylene and silicone-based coatings. Optionally in any embodiment, the coaxial cannula can be provided with a bevel cut profile, fillet cut or chamfer cut on its distal-most tip to ease the force of insertion through the bodies tissues.

Optionally in any embodiment, the hydrogel delivery needle and coaxial cannula can be provided with external graduation marks on their exterior surfaces to monitor the depth of insertion into tissue and also to determine the position of the coaxial cannula in relation to the delivery needle. These depth graduations can be created using laser marking or ink pad printing or similar. Spacing of 5-10 mm between graduation marks are typical.

Optionally in any embodiment, the methods described herein include an aspiration step to ensure no major blood vessel is punctured. This aspiration step may be conducted when the delivery needle is inserted into the target location and before the hydrogel plug is injected. This may be desirable so as to limit or prevent any hydrogel from entering into the vasculature which may result in a pulmonary embolism. Aspiration of dark blood would be an indication that a major blood vessel has been punctured.

Optionally in any embodiment, the hydrogel filled syringe employed can be configured to require aspiration before injection of the hydrogel material. To achieve this, a mechanism can be built into the syringe to restrict the forward actuation of the syringe plunger until a retracting aspiration actuation has been performed.

Optionally in any embodiment the system describe herein may include an additional empty syringe for the purpose of performing the aspiration step.

Optionally in any embodiment the device may contain a 2- or 3-way medical stopcock fluidically attached to the delivery device. Any or both of the hydrogel filled syringe and the aspiration syringe may be attached to the delivery device via the medical stopcock which can be actuated to change and restrict the fluid delivery path between aspiration syringe and hydrogel filled syringe. This may provide the advantage of allowing a faster aspiration and injection step and reduce the time spend in the lung prior to injection of the hydrogel plug.

Optionally in any embodiment, the syringe is an ergonomic syringe for improved deliverability. Examples are described in US20090093787 A1 ‘Ergonomic Syringe’ and U.S. Pat. No. 6,616,634 B2 ‘Ergonomic Syringe’. The system may also include an ergonomic syringe adapter which can be mounted onto the syringe. An example is described in USD675317 S1 ‘Ergonomic syringe adapter’. The syringe may include a mechanism to inject the viscoelastic hydrogel under high pressure. This may be in the form of a syringe assist device Optionally in any embodiment, the coaxial needle may have an internal sealing/valve feature that prevents any gel from entering the coaxial needle.

Optionally in any embodiment, the hydrogel delivery needle can be employed as a core needle within the coaxial needle.

Optionally in any embodiment, the positioning mechanism also comprises a firing mechanism, for example a spring-loaded firing mechanism, to quickly advance the delivery needle through the coaxial cannula to a predetermined depth. The required distance can either be a set distance for penetration depth, or can be adjustable to take into account the coaxial cannula position in relation to the target injection site. The device can be positioned using measurements taken through imaging.

The system, device and methods of the invention may employ a coaxial needle with a core that has a radiolucent marker for more accurate determination of position.

Optionally in any embodiment the delivery device can be provided in an elongated and flexible configuration so that it can be passed through an endoscope to perform injections at predetermined injection depths via an endoscope. The elongated members can include both the coaxial cannula and delivery needle elements of the delivery device.

Optionally in any embodiment the delivery device can be provided with one or multiple energy delivery elements that can deliver sufficient energy into a target location so as to bring about a therapeutic effect. The elements can be positioned at the distal-most tip of the needle, or proximal to the distal-most tip. The delivered energy can be in the form of electrical, radiofrequency, thermal (including heating and cooling effect), microwave, short wave or acoustic energy. The energy delivering device can be connected at its proximal end to a power source which can include control and feedback capabilities. Irrigation channels can be incorporated in the delivery device to provide coolant to the treatment site during treatment. A typical application of this treatment would include cancer ablation.

Optionally in any embodiment the delivery device can be provided with sensors to provide feedback as to the local and/or surrounding tissue parameters including electrical, chemical, optical, acoustic, mechanical and thermal. Sensors can be disposed proximate, distal to and proximal to the hydrogel outlet.

In another aspect, the invention provides a method of performing a lung procedure (for example a lung biopsy or a lung ablation procedure), comprising the steps of:

advancing a coaxial cannula into the lung, wherein a distal portion of the coaxial cannula has one or more apertures in a side wall thereof;

advancing a lung procedure needle through the cannula to a procedure site within the lung; actuating the lung procedure needle to perform a lung procedure at the procedure site;

withdrawing the lung procedure needle through the cannula;

advancing a hydrogel delivery needle through the coaxial cannula, wherein a distal portion of the hydrogel delivery needle has one or more apertures in a side wall thereof corresponding to the one or more apertures in the side wall of the coaxial cannula;

aligning the one or more apertures of the coaxial cannula and hydrogel delivery needle;

injecting a composite viscoelastic hydrogel of the invention through the one or more outlets in the hydrogel delivery needle and one or more outlets of the coaxial cannula into the lung to form a sealing plug that embraces the coaxial cannula and typically abuts the visceral pleura; and

withdrawing the coaxial cannula and hydrogel delivery needle through the sealing plug.

In one embodiment, the composite viscoelastic hydrogel is delivered adjacent the visceral pleura of the lung. In one embodiment, the lung procedure needle is a biopsy needle. In one embodiment, the lung procedure needle is a tissue ablation probe.

BRIEF DESCRIPTION OF THE FIGURES

FIGS. 1A-1D. Series of lateral views illustrating a transthoracic needle biopsy procedure and demonstrating how a pneumothorax occurs (prior art).

FIGS. 2A-2D. Series of lateral views illustrating embodiments of the delivery device and a method of delivering a hydrogel plug to a target location in the lung.

FIGS. 3A-3C. Series of graphs demonstrating the effect of crosslinked gelatin powder particle size on the dynamic viscoelastic properties of biphasic hydrogels.

FIGS. 4A-4C. Series of graphs demonstrating the influence of varying concentrations of crosslinked gelatin powder and hyaluronic acid on the dynamic viscoelastic properties of biphasic hydrogels.

FIG. 5. Graph showing the effect of varying concentrations of crosslinked gelatin powder and hyaluronic acid on the compressive modulus of biphasic hydrogels.

FIG. 6. Graph demonstrating the shear thinning behaviour of biphasic hydrogels.

FIGS. 7A-7D. Series of charts showing the effect of terminal steam sterilization on the dynamic viscoelastic properties of the biphasic hydrogel.

FIG. 8. Graph showing the effect of dehydrothermal treatment temperature on the viscosity of the biphasic hydrogel.

FIGS. 9A-9C. A section of a CT scan of the lung demonstrating the degradation of the biphasic hydrogel in vivo.

FIG. 10. Graph showing the injection force required to inject various biphasic hydrogels through an 18G delivery needle.

DETAILED DESCRIPTION OF THE INVENTION

All publications, patents, patent applications and other references mentioned herein are hereby incorporated by reference in their entirety for all purposes as if each individual publication, patent or patent application were specifically and individually indicated to be incorporated by reference and the content thereof recited in full.

The high efficacy demonstrated by exemplary embodiments disclosed herein is due to the unique viscoelastic properties of the hydrogel delivered. A hydrogel has both flow and elastic properties. Elasticity is reversible deformation; i.e. the deformed body recovers its original shape. The mechanical properties of an elastic solid may be studied by applying a stress and measuring the deformation of strain. Flow properties are defined by resistance to flow (i.e. viscosity) and can be measured by determining the resistance to flow when a fluid is sheared between two surfaces. The physical properties of a gel by viscoelasticity can be expressed by dynamic viscoelastic characteristics such as storage modulus (G′), loss modulus (G″), tangent delta (tan δ) and the like. Storage modulus characterizes the firmness of a composition and describes the storage of energy from the motion of the composition. Viscous modulus is also known as the loss modulus because it describes the energy that is lost as viscous dissipation. Tan δ is the ratio of the viscous modulus and the elastic modulus, tan δ=G″/G′. A high storage modulus and a low loss modulus indicate high elasticity, meaning a hard gel. Reversely, a high loss modulus and a low storage modulus mean a gel with high viscosity.

When the hydrogel described herein is used as a biomedical material, e.g., a biodegradable hydrogel plug for use in the periphery of the lung to prevent pneumothorax, it is considered that the increased stiffness and storage modulus of the gel can bring about improvement in sealing and barrier effect between tissues. It would also contribute to a prolonged duration (increased retention) at the target site, especially if the elasticity is greater than the elasticity of the surrounding tissues. The flowable nature of the hydrogel is due to its high Tan δ and at rest this allows for improvement in apposition with the surrounding tissue. This flow property also provides the hydrogel with its self-healing ability.

Therefore, it is preferably desirable that the gel for such use have well-balanced elasticity and viscosity. If the hydrogel zero shear viscosity is too high and if the gel does not portray sufficient shear thinning properties, it may become too difficult to inject through the delivery device into the target site. The gel may not readily appose surrounding tissue to form a barrier against fluid leak. Also, the gel may not readily flow back into the needle tract once the needle has been removed. On the other hand, if tan δ exceeds 0.8, the gel behaves like a solution, and it may infiltrate the surrounding tissue or be ejected from the needle tract. That is, the hydrogel described herein is regarded to have the most suitable physicochemical and rheological properties as a viscous plug for lung biopsy.

The term “viscoelastic hydrogel” therefore refers to a hydrogel that exhibits viscoelastic properties. It generally has a storage modulus (G′) of preferably greater than 400 Pa, more preferably greater than 800 Pa and even more preferably greater than 1000 Pa. The viscoelastic hydrogel may exhibit a tangent delta (tan δ; G″/G′) of from 0.05 to 0.8, preferably from 0.1 to 0.5 and more preferably from 0.2-0.5 in dynamic viscoelasticity at a frequency of 1 Hz. Preferably, the viscoelastic hydrogel exhibits a loss modulus (G″) of from 200 to 6000 Pa, more preferably from 400 to 2000 Pa, in dynamic viscoelasticity at a frequency of 1 Hz at 25° C. The viscoelastic hydrogel may be free of crosslinking, lightly crosslinked, or strongly crosslinked to provide appropriate characteristics, for example to increase its storage modulus (G′) or to increase its in vivo residence time.

As used herein, the term “shear thinning” as applied to a hydrogel means that when shear stress is applied to the hydrogel, the storage modulus (G′) reduces, the tan δ increases and the overall viscosity reduces. This property provides injectable properties to the hydrogel. And allows it to be injected through a narrow-gauge needle, such as used in minimally invasive procedures such as lung biopsy (17-20 gauge) or lung ablation (10-14 gauge). The shear thinning hydrogel described herein typically exhibits a range of a storage modulus (G′) of 1-100 Pa, preferably from 1-50 Pa in dynamic viscoelasticity at a frequency of 1 Hz and 100% strain. Furthermore, the hydrogel described herein has self-healing properties and retain their high storage modulus (G′) and loss modulus (G″) when the shear strain is removed.

The hydrogel described herein possess shear thinning capabilities. That is, when shear stress is applied, the storage modulus (G′) reduces, the tan δ increases and the overall viscosity reduces. This property allows the gels to be injected through a narrow-gauge needle, such as used in minimally invasive procedures such as lung biopsy. The gel described herein portrays the physical properties with ranges of a storage modulus (G′) of less than 100 Pa, preferably less than 50 Pa in dynamic viscoelasticity at a frequency of 1 Hz and 100% strain. Furthermore, the gels described herein portrays rapid thixotropic recovery properties and retain their high storage modulus (G′) and loss modulus (G″) immediately on removal of the high shear rate.

As used herein, the term “self-healing” as applied to a viscoelastic hydrogel of the invention refers to the ability of the hydrogel to reform together. “Self-healing” may also be described as the ability of the hydrogel to spontaneously form new bonds when old bonds are broken within the material. As an example, when an annular sealing plug of viscoelastic hydrogel is delivered around a delivery needle, a self-healing viscoelastic hydrogel will flow back together once the needle is removed to form a non-annular sealing plug, typically consisting of a single-bodied cohesive matrix.

Optionally in any embodiment the sealing hydrogel plug should be able to self-heal a channel through its centre independent of its in vivo environment. By this we refer to the ability of the hydrogel to fill the channel through a time dependent viscoelastic flow mechanism.

Optionally in any embodiment the sealing hydrogel plug should be able to self-heal a channel through its centre dependent on its in vivo environment. Stresses from the in vivo environment imposed on the hydrogel plug may improve its ability to self-heal in a shorter duration compared to an uninterrupted plug.

Optionally in any embodiment, the composite viscoelastic hydrogel should be able to self-heal under its own weight without any influence from the surrounding environment. This may be demonstrated by creating a singular mass of the hydrogel, for example a sphere of the hydrogel created using approximately 0.5 ml of hydrogel. A cylindrical channel can be created through the centre of the sphere by passing a 17 gauge needle through its centre and retracting the needle. The sphere with the cylindrical channel through its centre can be placed at rest on a bench with the axis of the cylindrical channel perpendicular to the bend. The size of the channel can be monitored over time. Referring to the viscoelastic hydrogels described in this invention, specifically hydrogels comprising 2-6% hyaluronic acid, the following are the observations: initially the channel in the ball will be visible, but over time (1-15 mins, depending on the hydrogel formulation) this channel will close over as the hydrogel self-heals. This is as a result of the time dependent flow of the hydrogel.

Optionally in any embodiment, part or all of the composite viscoelastic hydrogel is comprised of a hyaluronan hydrogel. The hyaluronan polymer forms a continuous phase throughout the three-dimensional matrix. Optionally in any embodiment, the viscoelastic hydrogel is a high molecular weight hyaluronan hydrogel. Optionally in any embodiment, the composite viscoelastic hydrogel is a shear thinning hydrogel (viscosity decreases under shear strain). Examples of polymer materials that may be employed to make a viscoelastic hydrogel include hyaluronan, especially high molecular weight hyaluronan. Other hydrogel materials suitable for use in the present invention are outlined in the review articles ‘Shear-thinning hydrogels for biomedical applications’, Soft Matter, (2012) 8, 260, ‘Injectable matrices and scaffolds for drug delivery in tissue engineering’ Adv Drug Deliv Rev (2007) 59, 263-272, and ‘Recent development and biomedical applications of self-healing hydrogels’ Expert Opin Drug Deliv (2017) 23: 1-15.

As used herein, the term “hyaluronan” or “hyaluronic acid” or “HA” refers to the anionic non-sulphated glycosaminoglycan that forms part of the extracellular matrix in humans and consists of a repeating disaccharide→4)-β-d-GlcpA-(1→3)-β-d-GlcpNAc-(1→, or any salt thereof. Hyaluronan is the conjugate base of hyaluronic acid, however the two terms are used interchangeably. When a salt of hyaluronic acid is employed, the salt is generally a sodium salt, although the salt may be employed such a calcium or potassium salts. The hyaluronic acid or hyaluronan may be obtained from any source, including bacterial sources. Hyaluronic acid sodium salt from Streptococcus equi is sold by Sigma-Aldrich under the product reference 53747-1G and 53747-10G. Microbial production of hyaluronic acid is described in Liu et al (Microb Cell Fact. 2011; 10:99). The term also includes derivatives of hyaluronic acid, for example hyaluronic acid derivatised with cationic groups as disclosed in US2009/0281056 and US2010/0197904, and other types of functionalised derivatives, such as the derivatives disclosed in Menaa et al (J. Biotechnol Biomaterial S3:001 (2011)), Schante et al (Carbohydrate Polymers 85 (2011)), EP0138572, EP0216453, EP1095064, EP0702699, EP0341745, EP1313772 and EP1339753.

Hyaluronic acid can be categorised according to its molecular weight. High molecular weight (preferably>1000 kDa (1 Mda)), medium molecular weight (preferably 250-1000 kDa), low molecular weight (preferably 10-250 kDa), and oligo hyaluronic acid (preferably<10 kDa). The effect of molecular weight on hyaluronic acid hydrogel viscosity has previously been reported. The stiffness and viscosity of the final gel is dependent on both molecular weight and solution concentration. In studying the rheological properties of hyaluronic acid with different molecular weights, Rheological and cohesive properties of hyaluronic acid J Biomed Mat Res, 76A, 4, Pg 721-728, Falcone et al found that high molecular weight hyaluronic acid is considerably more cohesive than low molecular weight hyaluronic acid. It has been shown that the presence of high molecular weight hyaluronic acid hydrogels at a wound site leads to reduction in scarring. High molecular weight hyaluronic acid has been shown to be anti-inflammatory, enhanced angiogenesis and enhanced immunosuppression. Jiang et al found that high molecular weight hyaluronic acid has been shown to protect from epithelial apoptosis in lung injury “Regulation of lung injury and repair by Toll-like receptors and hyaluronan” Nature Medicine (2005) 11, 11 1173-1179. Furthermore, inhalation of high molecular weight hyaluronic acid has been used to treat lung conditions such as bacterial rhinopharyngitis, chronic bronchitis, cystic fibrosis and asthma. In some embodiments, the hyaluronic acid compositions of the hydrogel are free from crosslinking and are free from other therapeutic agents. Hyaluronic acid based hydrogels with characteristics potentially suitable for this application are described in U.S. Pat. No. 9,492,474B2. ‘Compositions of’ hyaluronan with high elasticity and uses thereof. This document describes a material, Elastovisc™, comprised of high concentration and molecular weight hyaluronic acid. Its intended use is for injection into joints to relieve pain and treat osteoarthritis.

As used herein, the term “hyaluronan hydrogel” preferably includes a three-dimensional network of hyaluronan polymers in a water dispersion medium. The hyaluronan polymer forms a continuous phase throughout the three-dimensional matrix. Optionally in any embodiment, the hyaluronan polymers are non-crosslinked. Optionally in any embodiment, the hydrogel is free of a crosslinking agent. Optionally in any embodiment, the matrix is formed with a homopolymer, typically a hyaluronic acid homopolymer. Optionally in any embodiment, the hydrogel is pH balanced or buffered to match the pH of the physiological environment. Optionally in any embodiment, the matrix is lightly crosslinked. Any crosslinking agent known to crosslink hyaluronic acid may be used for this purpose. Crosslinking agents may include epichlorohydrin, divinyl sulfone, I, 4-bis (2,3-epoxypropoxy) butane (or I, 4-bis (glycidyloxy) butane or 1,4 butanediol diglycidyl ether=BDDE), the I, 2-bis (2,3-epoxypropoxy) ethylene, I-(2,3-epoxypropyl)-2, 3-epoxy cyclohexane.

Optionally in any embodiment, the continuous phase of the composite viscoelastic hydrogel may be comprised of ‘multi-component’ hydrogel which refers to at least two hydrogels that are evenly blended and dispersed together to form a homogenous hydrogel continuous phase. This construct may also be referred to as a semi-interpenetrating polymer (hydrogel) network or interpenetrating polymer (hydrogel) network comprised of two or more hydrogels. As an example, a hyaluronan hydrogel (concentration may range from 1-5%) may be blended with a methylcellulose hydrogel (concentration may range from 3-15%). In the same manner, more than two hydrogels may be combined to form a single cohesive network whereby each hydrogel provides improved properties to the overall network. The properties of each hydrogels may be provided to increase stiffness and viscosity, to provide improved injectability (shear thinning), to provide improved self-healing, to prolong the residence (biodegradation) time of the hydrogel in vivo, to provide haemostatic properties, to provide antibacterial properties, to provide anti-inflammatory properties, to provide anti-coagulant properties, to provide pro-coagulant properties, to provide colour and marking capability (under visible and radiographic detection), to provide some diagnostic or therapeutic effect (for example chemotherapy), to provide resistance to extremes of heat (hot and cold), to provide improved biocompatibility, and to improve manufacturability and preparation of the overall hydrogel. One or more of these hydrogels may be crosslinked to provide improved properties, for example to increase the residence time of the hydrogel in vivo

The viscoelastic hydrogel is generally a “biphasic” hydrogel, which refers to a hydrogel formed by combining (through mixing or blending) a colloidal hydrogel with a continuous phase hydrogel. The colloidal hydrogel will form an evenly dispersed phase in the continuous hydrogel phase. A variety of natural and synthetic biodegradable polymers can be used to form the continuous hydrogel phase. Glycosaminoglycans, for example hyaluronan and its derivatives form one example. The hyaluronan may be preferably non-crosslinked or possibly lightly crosslinked so as to retain its viscoelastic properties, especially its shear thinning and self-healing ability. Optionally in any embodiment, the hyaluronan may be provided at concentrations of 0.4-6%. A variety of biodegradable polymers are also suited to form the colloidal hydrogel phase as outlined previously (collagen and gelatin are two examples). The colloidal hydrogel phase can be added in sufficient quantities to provide the advantage of increased residence time of the hydrogel in vivo. This can allow the necessary time to provide for healing of the tissue. An additional benefit is that an increased residence time can provide a long-term marking function of the biopsy side for use under video-assisted thoracoscopic (VATS) surgery. A suitable polymer is one that is insoluble in an aqueous environment and can be achieved by crosslinking of the polymer through conventional means. An example would be dehydrothermally crosslinked gelatin. It should be noted that by introducing a too large amount of the colloidal hydrogel phase, it may jeopardize the injectability and self-healing ability of such compositions. Optionally in any embodiment, the “biphasic” hydrogel can comprise a colloidal hydrogel at concentrations of 0.2-30%, 15-28%, or 20-25% of hydrogel forming polymer (w/v). Optionally, in any embodiment, the ratio of continuous phase polymer to dispersed phase polymer in the viscoelastic hydrogel is about 1:10 to 1:40, more preferably about 1:10 to 1:20.

Optionally in any embodiment, the composite viscoelastic hydrogel exhibits a storage modulus (G′) of greater than 400 Pa, more preferably greater than 600 Pa, more preferably greater than 800 Pa, more preferably greater than 1000 Pa. Optionally in any embodiment, the viscoelastic hydrogel exhibits tan δ (G″/G′) from 0.01 to 0.8, more preferably 0.1 to 0.6 in dynamic viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate at 25° C.

Optionally in any embodiment, the composite viscoelastic hydrogel may be provided as a powder that is reconstituted in a physiologically acceptable fluid, for example water, saline, autologous blood, or autologous plasma prior to the surgical procedure. Synthetic fluids such as low molecular weight PEG and glycerol may also be employed. The powder may be comprised of any suitable biocompatible polymer or combinations of polymers. In one embodiment, the powder may be provided in the hydrogel delivery needle. In one embodiment, the powder may be provided in a syringe with a suitable reconstitution fluid provided in a second syringe. In one embodiment, the powder has an average particle size of 1-500, 10-100 or 30-40 microns. The powder may be both regular or irregular in both shape, morphology and size distribution and may be formed through milling or other means known in the art. In certain instances, powder hydration can be controlled by varying the level of de-hydration of the powder particles such as in the case of collagenous based materials, for example collagen or gelatin.

Optionally in any embodiment, the hydrogel described herein may be provided in separate components, for example in multiple syringes and the means can be provided to allow mixing of the components prior to injection through the syringe. Crosslinking agents can be provided in one or more of these components to provide the material characteristics necessary to achieve a shear thinning and self-healing hydrogel. Mixing can be achieved by reciprocating the contents between the syringes and a static mixer can be employed to speed up this process.

In any embodiment the composite viscoelastic hydrogel composition can be provided in a physiological buffer, e.g., a phosphate buffer or a bicarbonate buffer. In some embodiments, the pH of the composition is between pH 5 and pH 9 or between pH 7.5 and pH 8.5. In some embodiments, the pH of the composition is 8.0. In some embodiments, the pH of the composition is 7.5. In some embodiments, the pH of the composition is 8.5. If needed, acid (such as HCL) or base (such as NaOH) can be added to the composition to attain the desired pH. In a specific embodiment, the hyaluronic acid hydrogel described herein consists essentially of hyaluronic acid present at a concentration of 50 mg/ml (or about 5% W/V, and having an average molecular weight of between 1-2 Mda. Ranges intermediate to the recited values are also intended to be part of this invention. For example, hyaluronan content in the compositions described herein may be between about 0.5% and about 6% (weight/volume). It should further be appreciated that the amount of hyaluronan in a particular volume may also be expressed by alternative means (e.g., mg/ml, gram/litre or mol/litre). A person of ordinary skill in the art would know how to convert the various means of expressing the amount of hyaluronan in a particular volume

As used herein, the term “sealing plug”, “hydrogel plug” or “gel plug” refers to a single body of viscoelastic hydrogel, for example biphasic composite hydrogel, that is suitable for delivery through a needle to a locus in the lung and which has sufficient viscoelasticity to push away the tissue surrounding the needle and coalesce to form a single closed annular sealing plug around the needle. The viscoelastic properties and stiffness of the gel prevents infiltration of the tissue, allowing the gel to precisely oppose the tissue and form an effective seal around the needle and subsequently cannula thereby preventing air from lungs leaking past the plug. The viscoelastic behaviour of the hydrogel allows the annular plug to coalesce upon removal of the cannula filling the hole in the annular plug and bearing against the visceral pleura to seal it after withdrawal of the coaxial cannula.

Optionally in any embodiment, the hydrogel plug should exhibit “limited-swelling” behaviour which means that its bulk size should not increase by any profound extent when placed in vivo, for example below the surface of the lung to prevent pneumothorax. A hydrogel plug that swells by a significant degree may cause unwanted physiological or biological effect. Some swelling of hydrogels in vivo is to be expected but in order to preserve the native tissue, swelling of the hydrogel plug should be limited. Swelling can be characterised by forming a predetermined size of hydrogel sphere, for example rolling 500 μl of hydrogel into a sphere, and by placing this ball of hydrogel into an aqueous solution. This volume 500 μl will initially equate to a sphere with a diameter of approx. 10 mm. The aqueous solution may be a saline or simulated body fluid solution and it may also contain the correct enzyme activity that is found in vivo. The size and shape and dissolution of the ball of hydrogel can then be monitored over a prolonged period of time. The swelling ratio can be determine from:

Swelling (%)=(Ws−Wd)/Wd×100

[Wd=Weight of polymer; Ws=weight of swollen polymer]

Preferably the selling ratio should not exceed 250%, more preferably it should not exceed 150%, and more preferably it should not exceed 130%. Sample degradation can be determined by comparing the dry weight of the polymer over time. Dry weight can be determined by lyophilising the samples. The degradation rate can be inferred from the remaining weight of the hydrogel:

Remaining Hydrogel (%)=(W2−W1)/W1×100

[W1=Original dry weight of polymer; W2=time dependent dry weight of polymer]

Different polymeric materials with thermo-responsive, shear-thinning, shape memory and biological properties can be combined to yield composite hydrogels with improved properties for this application. Improvements can include enhanced biocompatibility, injectability, viscosity, altered biodegradation, drug attachment, tissue adhesion, cohesiveness, sealing ability stability, hydrophilicity. Gelatin and hyaluronic acid are two examples. Substances which can be combined with these polymer include methylcellulose, oxidized cellulose, carboxylmethyl cellulose, and carboxylic acid.

Optionally in any embodiment, the composite viscoelastic hydrogel can include contrast medium which refers to an additive that can be included in the gel in an appropriate amount that allows the hydrogel to be contrasted against the surrounding tissue. In this way, the hydrogel plug and injected location can be visually identified and/or targeted for example during the surgical procedure or during a follow up surgical procedure.

Identification can be visual or through guidance systems such as CT scans, ultrasound or fluoroscopy. Additives which can be added to the hydrogel in varying concentrations to achieve effective visual contrast include ionic and non-ionic contrast medium, methylene blue, indigo carmine, toluidine blue, lymphazurine, hemotoxylin, eosin, indocyanine green (ICG), India ink, carbon based powders such as carbon black, carbon nanotubes and graphene, and ceramic powders such as aluminium oxide, titanium dioxide, and calcium phosphates. The hydrogel may also comprise additional detectable marking agents. The detectable marking agent suitable for use in the hydrogel described herein may include any composition detectable by spectroscopic, photochemical, biochemical, immunochemical, electrical, optical or chemical means. A wide variety of appropriate detectable markers are known in the art, which include luminescent labels, radioactive isotope labels, and enzymatic labels. These marking agents can be mixed with the hydrogel or chemically conjugated to the hydrogel molecules.

Optionally in any embodiment, the composite viscoelastic hydrogel can comprise of a therapeutic agent or biologically active agent. Therapeutic agents which may be linked to, or embedded in, the hydrogel include, but are not limited to, analgesics, anaesthetics, antifungals, antibiotics, anti-inflammatories, anthelmintics, antidotes, antiemetics, antihistamines, antihypertensives, antimalarials, antimicrobials, antioxidants, antipsychotics, antipyretics, antiseptics, antiarthritics, antituberculotics, antitussives, antivirals, cardioactive drugs, cathartics, chemotherapeutic agents, a colored or fluorescent imaging agent, corticoids (such as steroids), antidepressants, depressants, diagnostic aids, diuretics, enzymes, expectorants, hormones, hypnotics, minerals, nutritional supplements, parasympathomimetics, potassium supplements, radiation sensitizers, a radioisotope, sedatives, stimulants, sympathomimetics, tranquilizers, urinary anti-infectives, vasoconstrictors, vasodilators, vitamins, xanthine derivatives, and the like. Optionally in any embodiment, the hydrogel described herein comprises one or more anesthetics. Exemplary anesthetics include, but are not limited to, proparacaine, cocaine, procaine, tetracaine, hexylcaine, bupivacaine, lidocaine, benoxinate, mepivacaine, prilocalne, mexiletene, vadocaine and etidocaine. Optionally in any embodiment, the viscoelastic hydrogel can further comprise foaming agents, foam stabilizers, surfactants, thickeners, diluents, lubricants, wetting agents, plasticizers.

Optionally in any embodiment, part or all of the composite viscoelastic hydrogel can be “biodegradable” and configured to degrade over time in-vivo. Different phases or components of the viscoelastic hydrogel can be configured to degrade at different rates. Biodegradable substances are preferably eliminated by the body without causing an inflammatory or immune response. For the viscoelastic hydrogel described herein, the period of time for full biodegradation can be less than 1 year, preferably less than one month, more preferably less than 1 week, and more preferably less than 72 hours. The added benefit of a quick degradation period is that it allows the lung tissue to return to normal and prevents excess scar tissue formation at the delivery site. Also, limiting residence time and scar tissue formation ensures that the delivery of the hydrogel plug does not interfere with follow up radiological analysis of the suspected lung lesion. Non-crosslinked systems may result in a faster in vivo residence period compared to crosslinked systems. The high molecular weight (>1000 kDa) and high concentration (40-60 mg/ml) hyaluronic acid hydrogels described herein have a degradation period of less than 1 week and also less than 72 hours. Longer degradation periods are possible by modifying the native hyaluronic acid molecular structure via crosslinking or by other means. Longer degradation periods are also possible by combining the hyaluronic acid hydrogel with one or more hydrogels or colloidal hydrogels to form a composite hydrogel. One of the hydrogels will remain at the target site for a longer period while the other is removed. For example, the hyaluronic acid hydrogel may be combined with a crosslinked polymer (for example hyaluronan, hylan, collagen or gelatin) to form a composite hydrogel. The cross-linked polymer can be configured to have a residence time of greater than 1 week, and often greater than 2 weeks by the use of various crosslinking modalities known in the art. Cross-linkers employed as part of the implantable material precursors can include aldehydes, polyaldehydes, esters, and other chemical functionality suitable for cross-linking protein(s). Physical crosslinking methods can also be employed, for example subjecting the polymers to heat, cold or radiation. Crosslinking agents can be added to improve cohesion, rigidity, mechanical strength and barrier properties.

As used herein, the term “in-vivo residence time” as applied to a sealing plug of composite viscoelastic hydrogel refers to the period of time that sealing plug of 0.1-1 ml, preferably 0.2-0.8 ml and more preferably 0.3-0.5 ml that persists in lung tissue in-vivo without any significant loss of structure integrity. The in-vivo residence time should be sufficient to allow healing of the hole in the visceral pleura to occur, and ideally to allow for healing in the surrounding lung tissue to occur. Methods of approximating the in-vivo residence time of hydrogels are described below. To achieve an appropriate in-vivo residence time to allow healing to occur, the hydrogel can be comprised of certain unmodified materials (including proteins) that have a longer residence time. Examples include collagen, oxidised cellulose, starch, extracellular matrix (ECM). Crosslinked hydrogels as described herein have been found to have an in-vivo residence time of more than two weeks. Optionally, the shear-thinning viscoelastic hydrogel may have an in-vivo residence time of at least 1 week, preferably at least 2 weeks, and ideally at least 3 weeks.

EXEMPLIFICATION

The invention will now be described with reference to specific examples. These are merely exemplary and for illustrative purposes only: they are not intended to be limiting in any way to the scope of the monopoly claimed or to the invention described. These examples constitute the best mode currently contemplated for practicing the invention.

The mechanism of pneumothorax resulting from a transthoracic needle biopsy is illustrated in FIGS. 1A-1D (Prior art). FIG. 1A illustrates a cross section of the thoracic cavity A, which comprises the thoracic (chest) wall muscle B, ribs C, lung tissue D, and the pleural cavity E defined by the serous membrane of the thoracic wall (parietal pleura F) and the serous membrane of the lung (visceral pleura G). During a lung biopsy procedure (FIG. 1B), a core needle H and coaxial cannula I are advanced percutaneously through the skin O and through the pleural cavity E towards a suspected lung nodule J. In FIG. 1C the core needle H has been withdrawn and replaced with a biopsy needle K which is advanced through the cannula 2 and obtains a tissue sample from the suspected lung nodule J. As illustrated in FIG. 1D, removal of the biopsy needle K and cannula I leaves a void L in the lung tissue D and also leaves a hole L1 in the visceral pleura G. The dense muscular tissue of the thoracic wall B contracts around the void caused by removal of the needles. However, the holes L, L1 created by the biopsy needles in the lung tissue D and visceral pleura G do not completely seal over. Due to the pressure gradient between the lung tissue D and pleural cavity E, air escapes through the hole L1 created in the visceral pleura G and enters the pleural cavity E, creating a collection of air in the pleural cavity E known as a pneumothorax M. If a blood vessel of significant size is punctured during the biopsy procedure the pleural cavity may also fill with blood, a condition known as a haemothorax. The prevalence of haemothorax is not as high as pneumothorax. The haemothorax or pneumothorax M can grow in sufficient size to cause the lung to partially or fully collapse and bring about respiratory distress and the need for treatment.

Referring to FIGS. 2A-2E, a method for overcoming the shortcomings of the prior art is presented. In FIGS. 2A-2E, a method of delivering a viscoelastic hydrogel plug to a target location in the lung is described. This embodiment, employs a medical device system comprising a coaxial cannula 2 having a distal-most end 2A and a proximal connector such as a luer lock 2B, a core needle 3, and a hydrogel delivery needle 4 having a distal tissue piercing tip 5 and a hydrogel outlet 6 disposed on a side of the needle proximal of the piercing tip 5. Also contained in the system is a syringe 15 with reservoir 15 b filled with viscoelastic hydrogel material including any of those described herein. The syringe may be replaced by any pump, plunger, fluid advancement mechanism or element suitable for delivering a viscous hydrogel.

As shown in FIG. 2A, the core needle 3 and cannula 2 assembly are inserted into the chest wall of the patient to a depth at which the assembly is located in the chest wall B and does not penetrate the lung D. A coaxial cannula 2 refers to a needle device having an inner lumen configured to receive a penetrating device, for example a core needle 3 where the assembled core needle and cannula 2 may be used to enter through the skin surface on the chest. Generally, the coaxial cannula has a gauge size of 10 to 19. In additional embodiments, the coaxial cannula may also be referred to as a sheath, an introducer, an obturator/stylet assembly, a guiding catheter, trocar, port device or other medical introductory device known in the art.

As shown in FIG. 2B, the core needle 3 has been withdrawn from the cannula 2 and a hydrogel delivery needle 4 is advanced through the cannula 2. The hydrogel delivery needle 4 typically has a piercing tip, and a hydrogel outlet 6 which is typically disposed on a side of the needle proximal of the piercing tip 5, for example 0.5-15 mm from the piercing tip 5. The delivery needle 4 has a distal-most end configured for insertion into the body, and a proximal end which during use is positioned outside of the body. The needle is generally formed from a metal, although the positioning (adjustment) mechanism may be formed from plastic or polymer or a metal. The needle may comprise polymer tubing at its proximal end and may include a luer lock to facilitate fluidically connecting the needle (or polymer tubing part) to a pump or syringe 15. Generally, the hydrogel delivery needle 4 has a gauge of 13 to 20. The hydrogel delivery needle 4 is inserted to a depth at which the hydrogel outlet 6 is positioned in the lung tissue distal of the pleural cavity E and visceral pleura G. Positioning of the hydrogel outlet 5 at this target location may be achieved under CT guidance by employing a radiopaque or radiolucent marker 32 on the delivery needle which can be positioned a known distance X from the hydrogel outlet 6. By overlaying the radiolucent marker 32 with the pleural cavity E, the hydrogel outlet can be positioned a predetermined distance X inside the lung D from the pleural cavity E. The pleural cavity E is a very thin space approximately 25 μm in width and is often referred to as a virtual cavity. As can be seen later in FIG. 7, the pleural cavity E can be distinguished under CT guidance as the transition between the lung (dark area) and chest wall (bright area). Positioning of the radiolucent marker 32 over the pleural cavity E may be achieved by stepwise scanning and fine adjustment of the needle 4, or with fine adjustment under continuous fluoroscopic guidance.

As shown in FIG. 2C, a syringe 15 with hydrogel filled reservoir 15B is attached to the delivery needle 4 via a luer lock 12. A predefined quantity of viscoelastic hydrogel is then injected into the lung through the hydrogel outlet 6 to form a closed annular viscoelastic sealing plug 7 around the delivery needle 4. Subsequent to this step, the coaxial cannula 2 is advanced over the delivery needle 4 through the sealing plug 7 and towards the suspected lung nodule J. The hydrogel delivery needle 4 is withdrawn leaving the cannula 2 with surrounding hydrogel sealing plug 7 in place for receipt of a lung biopsy needle K. As shown in FIG. 2D a lung biopsy needle K can be then advanced through the cannula 2 and a lung biopsy carried out, The biopsy needle K and cannula 2 are both withdrawn after the biopsy has been taken. As shown in FIG. 2E the sealing plug 7 remains in position in the lung tissue after the needles have been withdrawn. Due to the physical properties of the viscoelastic hydrogel material, the sealing plug 7 reflows into the space left behind by the needles, as well as sealing the hole L1 left in the visceral pleura G by the coaxial cannula 2. These steps describe a method of performing a lung biopsy with diminished chance of causing a pneumothorax. The efficacy of the sealing plug 7 is dependent on its ability to block any air in the aerated lung tissue D from exiting the hole L1 in the visceral pleura G.

For a number of reasons it may be difficult to position the delivery device as outlined above. Firstly, fluoroscopic guidance may not be available to the clinician so that the delivery needle 4 with marker band 32 cannot be accurately positioned. Secondly, it may be harmful to expose the patient to too many CT scans and resulting high radiation dose to achieve accurate placement of the needle marker band 32. Furthermore, delayed placement of the hydrogel plug may lead to potential pneumothorax while the needle is in the lung tissue unprotected. In order to quickly, easily and accurately target the required depth of injection in the lung for the viscoelastic hydrogel to achieve an effective seal, a positioning mechanism is provided with the hydrogel delivery needle 4 as will be described hereafter.

EXAMPLES

Example 1: A composite viscoelastic hydrogel comprising hyaluronic acid and crosslinked gelatin was created using the following method. Type A porcine derived gelatin (300Bloom) was dissolved fully in water at 7% w/v at 40° C. and allowed to set at 4° C. overnight. The resulting gel were subsequently freeze dried by freezing at −40° C. and drying at 25° C. under a constant vacuum of 0.1 mbar. The dried constructs were then heated under vacuum conditions (0.001 mbar) for 24 hours at 140° C. to induce crosslinking. The sponge was then roughly diced before being milled to form a fine powder using a cryo-mill (Model: 75 Spex SamplePrep, LLC.). The powder was sieved using a 125 μm sieve and the resultant powders had a powder particle size distribution of D×10=7.4 μm, D×50=32.8 μm, D×90=95 μm as measured using a Mastersizer 3000 laser diffraction particle size analyser (Malvern Panalyticlal ltd). The dehydrothermally crosslinked gelatin powder was mixed with sodium hyaluronate powder (molecular weight: 1.8-2 MDa) and the powder mixture was hydrated with phosphate buffered saline solution at the following concentration: Gelatin: 130 mg/ml, Sodium hyaluronate: 35 mg/ml. The resulting hydrogel was loaded into a syringe. The hydrogel was employed to prevent pneumothorax during a CT-guided transthoracic needle biopsy procedure as outlined in FIG. 8A-8F. This procedure was performed in a porcine model. The hydrogel formed an annular sealing plug around the needle during the biopsy procedure and after the needles were withdrawn, the hydrogel self-healed to prevent pneumothorax. The hydrogel persisted at the site for at least 1 week as was evidence from CT-scan follow-up.

Example 2: A composite viscoelastic hydrogel comprising hyaluronic acid and crosslinked gelatin was created using the following method. A type A porcine derived gelatin powder (300bloom) was ground to a fine powder using a cryo-mill (Model: 75 Spex SamplePrep, LLC.). The powder was sieved using a 125 μm sieve and the resultant powders had a powder particle size distribution of D×10=5.4 μm, D×50=35.5 μm, D×90=90 μm as measured using a Mastersizer 3000 laser diffraction particle size analyser (Malvern Panalyticlal ltd). The resultant fine powder was heat treated under vacuum conditions (0.001 mbar) for 24 hours at 160° C. to induce crosslinking. The DHT crosslinked gelatin powder was mixed with sodium hyaluronate powder (molecular weight: 1.8-2 MDa) and the powder mixture was hydrated with phosphate buffered saline solution at the following concentration: Gelatin: 100 mg/ml, Sodium hyaluronate: 45 mg/ml. The resulting hydrogel was loaded into a syringe. The hydrogel was employed to prevent pneumothorax during a CT-guided transthoracic needle biopsy procedure similar to that outlined in FIG. 8A-8F. This procedure was performed in a porcine model. The hydrogel formed an annular sealing plug around the needle during the biopsy procedure and after the needles were withdrawn, the hydrogel self-healed to prevent pneumothorax.

Using the above method, various concentrations of the biphasic gel were evaluated rheologically and experimentally. The measurement of the dynamic viscoelasticity and dynamic viscosity of the hydrogels was made using a rheometer Model AR2000 manufactured by TA Instruments under the following conditions.

Method of measurement: oscillation test method, strain control

Measuring temperature: 25° C.

Geometry: 4° cone plate angle

Measuring geometry: 4 cm

Truncation gap: 112 μm

Frequency: 1 Hz

Storage Crosslinked Sodium Modulus @ Tanδ @ Gelatin Hyaluronate 1 Hz & 1 Hz & Zero shear Viscosity @ Concentration Concentration 1% Strain 1% Strain viscosity 100 s⁻¹ 100 mg/ml 45 mg/ml 5,813 Pa 0.4 18,367 Pa · s 6.8 Pa · s 150 mg/ml 45 mg/ml 11,667 Pa  0.27 43,317 Pa · s 10.0 Pa · s  100 mg/ml 35 mg/ml 2,722 Pa 0.45  6,700 Pa · s 4.2 Pa · s 150 mg/ml 35 mg/ml 6,406 Pa 0.37 14,150 Pa · s 5.9 Pa · s

Example 3: FIGS. 3A-3C. represents viscoelastic properties of composite viscoelastic hydrogels of the invention as measured using a dynamic oscillatory test method. Various biphasic viscoelastic hydrogels comprising hyaluronic acid and crosslinked gelatin powder were fabricated using the following method. A type A porcine derived gelatin powder (300bloom) was ground to a fine powder using a cryo-mill (Model: 75 Spex SamplePrep, LLC.). The powder was sieved using sieves of varying mesh size to yield powder particles of the following size ranges: <53 μm, 50-100 μm, 100-200 μm, 200-300 μm. The resultant gelatin powders were heat treated under vacuum conditions (0.001 mbar) for 24 hours at 140° C. to induce dehydrothermal (DHT) crosslinking. The DHT crosslinked gelatin powders were subsequently mixed with 7.5 mg/ml sodium hyaluronate powder (molecular weight: 1.8-2 MDa) at the following concentrations: 120 mg/ml, 140 mg/ml, 160 mg/ml. The powder mixtures were hydrated with phosphate buffered saline to form the biphasic hydrogels. The test rheometer used to measure the rheological properties of the hydrogels was a model MCR102 by Anton Paar GmbH. Dynamic oscillatory tests were conducted under stress control, with a 25 mm flat plate geometry, a gap of 1 mm, an analysis temperature of 25° C. and over the frequency range 0.1-10 Hz. The graphs present the Storage Modulus (G′) (FIG. 3A), Loss Modulus (FIG. 3B), and Tan Delta (G′/G″) (FIG. 3C) for the biphasic hydrogels taken at a frequency of 1 Hz. The data demonstrates how the Storage (G′) and Loss (G″) moduli both increase for increasing powder particle size, however the Tan δ (G′/G″) decreases for increasing powder particle size.

Example 4: FIGS. 4A-4C. represents viscoelastic properties of composite viscoelastic hydrogels of the invention as measured using a dynamic oscillatory test method as outlined above. Various biphasic viscoelastic hydrogels comprising hyaluronic acid and crosslinked gelatin powder were fabricated using the method outlined above. A uniform gelatin powder particle size (Pm) of 125 μm<Pm<300 μm was used to fabricate the hydrogels. The following hyaluronic acid concentrations were employed: 5 mg/ml, 7.5 mg/ml, 10 mg/ml, 15 mg/ml, 20 mg/ml. Each concentration was combined with one of the following gelatin concentrations: 120 mg/ml, 140 mg/ml, 160 mg/ml, 180 mg/ml to create a matrix of 20 biphasic hydrogels for rheological analysis. The graphs present the Storage Modulus (G′) (FIG. 4A), Loss Modulus (FIG. 4B), and Tan Delta (G′/G″) (FIG. 4C) for the biphasic hydrogels taken at a frequency of 1 Hz. The general trends are an increase in Storage (G′) and Loss (G″) moduli for increasing concentrations of both gelatin and hyaluronic acid. We also see a trend towards higher Tan δ (G′/G″) for increasing hyaluronic acid concentration.

Example 5: FIG. 5 presents results of compression testing to determine the compression stiffness of the composite viscoelastic hydrogels of the invention. Biphasic hydrogels, formed in a phosphate buffered saline carrier of varying hyaluronic acid and gelatin concentrations were prepared as previously described. Hyaluronic acid with a molecular weight of 1.8-2 MDa was employed.

Hydrogels were formed into 4 mm thick sheet and using a core biopsy punch, 6 mm diameter cylinders with a height of 5 mm were created. Compression tests of the cylindrical samples of hydrogel and lung tissue were performed using a Zwick universal testing machine with a 5N load cell at a strain rate of 3 mm/min. The results show an increase in compressive stiffness with increasing concentrations of gelatin and hyaluronic acid. A large spike in compressive stiffness is seen at DHT gelatin concentrations in excess of 140 mg/ml. Similar tests were performed with samples of parenchymal lung tissue that yielded a compressive stiffness of approx. 800 Pa.

Example 6: FIG. 6 shows the strain sweep data for composite viscoelastic hydrogels of the invention as measured using a dynamic oscillatory test. Biphasic hydrogels, formed in a phosphate buffered saline carrier of varying hyaluronic acid and gelatin concentrations were prepared as previously described. The test rheometer used to measure the rheological properties of the hydrogels was a model MCR102 by Anton Paar GmbH. Tests were conducted under strain control, with a 25 mm flat plate geometry, a gap of 1 mm, an analysis temperature of 25° C., a frequency of 1 Hz and over a strain range of 0.01-100%. The graph demonstrates how all gels exhibit shear thinning behaviour and all gels demonstrate a storage modulus G′ of less than 200 Pa at 100% strain.

Example 7: FIG. 7 demonstrates the effect of terminal steam sterilization on the dynamic viscoelastic properties of the composite viscoelastic hydrogel of the invention. Biphasic hydrogels were formulated consisting of varying concentrations of dehydrothermally crosslinked gelatin: 160 mg/ml, 170 mg/ml, 200 mg/ml. The gelatin was combined with hyaluronic acid in phosphate buffered saline as previously described. The gels were loaded into 1 ml glass syringes and steam autoclaved at 128° C. for 15 mins. The rheological characteristics of the hydrogels were determined both pre and post-sterilization The test rheometer used to measure the rheological properties of the hydrogels was a model MCR102 by Anton Paar GmbH. Dynamic oscillatory tests were conducted under stress control, with a 25 mm flat plate geometry, a gap of 1 mm, an analysis temperature of 25° C. and over the frequency range 0.1-10 Hz. The charts present the Storage Modulus (G′) (FIG. 7A), Loss Modulus (FIG. 7B), and Tan Delta (G′/G″) (FIG. 7C) for the biphasic hydrogels taken at a frequency of 1 Hz before and after sterilization. FIG. 7D represents the yield stress data. Yield stress was measured by exposing a sample to a steady stress ramp until the sample began to undergo plastic deformation; the stress at which plastic deformation begins was taken to be the yield stress. The shear rate was ramped logarithmically and had an initial value 0.0011/s and final value of 2001/s.

Example 8: FIG. 8 presents viscosity data for composite viscoelastic hydrogels of the invention produced using different dehydrothermal treatment conditions. Milled gelatin powder was heated under vacuum conditions (0.001 mbar) for 24 hours at three different temperatures (130° C., 140° C. and 150° C.) to induce dehydrothermal (DHT) crosslinking. This resulted in different levels of crosslinking for the powders, with increased temperature leading to a higher degree of crosslinking. Equal concentrations of gelatin powder from each batch were combined with hyaluronic acid (8 mg/ml) in phosphate buffered saline. Hydrogels were then evaluated rheologically. The test rheometer used to measure the rheological properties of the hydrogels was a model MCR102 by Anton Paar GmbH. Dynamic oscillatory tests were conducted under shear rate control, with a 25 mm flat plate geometry, a gap of 1 mm, an analysis temperature of 25° C. The viscosity was recorded at a shear rate of 0.0011/s. The results show a reduction in viscosity with increasing dehydrothermal treatment. This is likely owning to the fact that the higher level of crosslinking reduces the ability of the hydrogel particles to swell, therefore more residual fluid is left in the sample.

Example 9: FIG. 9A to FIG. 9C presents partial CT-scans of the cross-sections of a composite viscoelastic hydrogel plug used to prevent pneumothorax during transthoracic lung biopsy procedure in a porcine model. The bi-phasic hydrogel is formed by combining hyaluronic acid and dehydrothermally crosslinked gelatin as described previously. The procedure used to deliver the biphasic hydrogel plug is similar to the one described in FIG. 2A to FIG. 2E. At the start of the biopsy procedure a hydrogel plug of volume 0.3 ml is injected to just below the surface (visceral pleura) of the lung. The hydrogel material is tissue opposing and does not infiltrate the lung tissue. It therefore creates a plug surrounding the delivery needle and coaxial cannula. The hydrogel plug thereby creates an air-tight seal between the cannula and the surface of the lung. After the biopsy procedure the needles are removed, the biphasic hydrogel self-heals by flowing back to occupy the space taken up by the needle. This further prevents any air from escaping the lung. FIG. 9A shows a partial CT scan of the lung and in particular a cross-section of the hydrogel plug immediately after the lung biopsy procedure. The black arrow indicates the hydrogel plug. FIG. 9B shows a partial CT scan of the lung and the same hydrogel plug after 13 days. FIG. 9C shows a partial CT scan of the lung at the same location of the hydrogel plug after 30 days. It is event in FIG. 9C that the hydrogel plug has fully degraded leaving healthy lung tissue in it's place.

Example 10: FIG. 10 represents the injection force employed to delivery various composite hydrogels of the invention through an 18G delivery needle of 150 mm in length. The hydrogels were prepared as described previously with varying concentrations of DHT treated gelatin powder and 8 mg/ml hyaluronic acid in phosphate buffered saline. To perform the test, hydrogels were loaded into 1 ml glass syringes and the syringes were mounted onto a Zwick universal testing machine with a 200N load cell. An 18G delivery needle of length 150 mm was attached to the 1 ml glass syringe via a luer lock connector. A compression test was performed at a rate of 120 mm/min so that the load cell depressed the syringe plunger and the hydrogel was injected through the delivery needle. Results show that the maximum injection force achieved for all gels lay between 10-25N. This is believed to be below the clinically acceptable level of 45N.

Example 10: In order to determine the DHT gelatin powder particle size a laser diffraction technique was employed (Malvern Instruments Master Sizer 3000) using the following methods. Data was analysed using the respective software (Mastersizer 3000, v.0.1). An even dispersion of dry particles was achieved using the Aero S dispersion unit. The refractive index of gelatin was set to 1.543 while the absorption index was set to 0.010. Data was recorded when obscuration levels were between 0.10% and 10.00%. An even dispersion of wet particles was achieved using the Hydro LV dispersion unit. The refractive index of PBS and gelatin was set to 1.33 and 1.543, respectively. The absorption index of the dispersed gelatin particles was set to 0.01. For measurement of wet particle size distribution, gelatin was hydrated at room temperature for 24 hours prior to testing. Particles were added to approximately 125 mL of PBS in the dispersion unit until an obscuration of approximately 10% was obtained.

DHT particles were prepared by firstly grinding 300bloom type A porcine gelatin powder. This powder was sieved with a mesh size of 1125 μm and subjected to a DHT treatment of 140° C. for 24 hours. Laser diffraction revealed the following dry particle size distribution: 10^(th) (Dv(10)), 50^(th) (Dv(50)), and 90th (Dv(90)) percentiles of 19, 58, and 123 microns respectively. Gelatin particles were then hydrated for 24 hours prior and wet particle size distribution revealed the following results: Dv(10), Dv(50) and Dv(90) was 27, 107 and 229 microns respectively. This indicates a volume swelling factor of approximately 6.2 for the particles post hydration.

In a preferred embodiment, the composite viscoelastic hydrogel is capable of preventing pneumothorax during procedures requiring transthoracic needle access by being injected just below the visceral pleura of the lung and by having the following properties:

-   -   1. The hydrogel has low enough viscosity under shear stress         exerted by the syringe to enable the hydrogel to be injected to         the target site through a needle, catheter or other luminal         device.     -   2. Once exiting the needle the hydrogel undergoes a rapid         thixotropic recovery to a stiffness sufficient to prevent         infiltration of lung tissue.     -   3. Once the needle has been removed, an element of viscous flow         enables the gel to flow back to form a single entity. The gel         flows back to fill the void left by the needle in the lung         tissue and in the visceral pleura. It may achieve this by having         a sufficient flowable nature which is preferably dependent on         having a high tan δ.     -   4. The gel has sufficient rigidity and storage modulus (G′) that         it is not prematurely ejected from the lung and remains at the         delivery site until healing has occurred.

EQUIVALENTS

The foregoing description details presently preferred embodiments of the present invention. Numerous modifications and variations in practice thereof are expected to occur to those skilled in the art upon consideration of these descriptions. Those modifications and variations are intended to be encompassed within the claims appended hereto. 

1. A composite viscoelastic hydrogel comprising: a continuous phase of non-crosslinked biodegradable hyaluronic acid gel; and a dispersed phase of dehydrothermally-crosslinked micron-sized gelatin hydrogel particles.
 2. A composite viscoelastic hydrogel according to claim 1 that exhibits a storage modulus (G′) of greater than 400 Pa and a tan δ (G″/G′) from 0.1 to 0.8 in dynamic viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate at 25° C.
 3. A composite viscoelastic hydrogel according to claim 1 that exhibits an axial compressive stiffness of greater than 800 Pa, as measured using an axial compression testing machine.
 4. A composite viscoelastic hydrogel according to claim 1, in which the dehydrothermally crosslinked micron-sized gelatin hydrogel particles have an average dimension of less than 100 microns prior to hydration.
 5. A composite viscoelastic hydrogel according to claim 1, that possesses an in vivo degradation period in the lung tissue of at least 2 weeks.
 6. A composite viscoelastic hydrogel according to claim 1, in which the dehydrothermally-crosslinked micron-sized gelatin hydrogel particles have an in-vivo degradation period in the lung tissue of less than 2 months.
 7. A composite viscoelastic hydrogel according to claim 1 which is shear-thinning.
 8. A composite viscoelastic hydrogel according to claim 1, that is shear-thinning and demonstrates a storage modulus G′ of less than 200 Pa at 100% strain as measured by a rheometer under strain control and at a test frequency of 1 Hz.
 9. A composite viscoelastic hydrogel according to claim 1, comprising 8-25% gelatin (w/v).
 10. A composite viscoelastic hydrogel according to claim 1, comprising about 0.4-6% of hyaluronic acid (w/v).
 11. A composite viscoelastic hydrogel according to claim 1, whereby the hydrogel is terminally sterilized via steam sterilization.
 12. A composite viscoelastic shear-thinning hydrogel according to claim 1, comprising: a continuous phase of non-crosslinked biodegradable hyaluronic gel; and a dispersed phase of dehydrothermally-crosslinked micron-sized gelatin hydrogel particles, in which the hydrogel exhibits a storage modulus (G′) of greater than 400 Pa and a tan δ (G″/G′) from 0.01 to 0.8 in dynamic viscoelasticity measured by a rheometer at 1 Hz and 1% strain rate at 25° C.
 13. A composite viscoelastic shear-thinning hydrogel according to claim 1, comprising: a continuous phase of non-crosslinked biodegradable hyaluronic gel; and a dispersed phase of dehydrothermally-crosslinked micron-sized gelatin hydrogel particles, in which the hydrogel comprises 8-25% gelatin hydrogel particles (w/v) having an average dimension of less than 100 microns in a dehydrated state, and about 0.5-2.0% of hyaluronic acid (w/v).
 14. A composite viscoelastic shear-thinning hydrogel according to claim 1, comprising: a continuous phase of non-crosslinked biodegradable hyaluronic gel; and a dispersed phase of dehydrothermally-crosslinked micron-sized gelatin or collagen hydrogel particles, in which the hydrogel exhibits an in vivo degradation period in the lung tissue of at least 2 weeks, and in which dehydrothermally-crosslinked micron-sized gelatin hydrogel particles exhibit an in-vivo degradation period in the lung tissue of less than 2 months.
 15. (canceled)
 16. (canceled)
 17. (canceled)
 18. (canceled) 